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Future Perspectives\nResearch and development in science and technology should not be discrete. If we evaluate correctly and examine carefully what have been done in the past, we will be able to foresee what would be available for us in future. Titanium materials science is a typical interdisciplinary science and engineering. In order for engineered titanium materials to serve as titanium biomaterials, we have been discussing and reviewing numerous articles to prove that appropriate surface modifications and characterizations should be properly preformed and reflected to appropriate fabrication technologies and methods.\nIn the past, when tissues became diseased or damaged, a physician had little recourse but to remove the offending part, with obvious limitations. Removal of joints, vertebrae, teeth, or organs led to only a marginally improved quality of life. However, human survivability seldom exceeded the progressive decrease in quality of tissues, so the need for replacement parts was small. During the last century the situation changed greatly. The discovery of antiseptics, penicillin and other antibiotics, chemical treatment of water supplies, improved hygiene, and vaccination all contributed to a major increase in human survivability in developed countries. Life expectancy is now in the range of 80+ years. In the past, it was the major practice to remove the diseased tissues, whereas at the present, either transplants (using autografts, heterografts, or homografts) or implants (using biological fixation, bioactive fixation, or cement fixation) are commonly utilized, and in the future, regeneration of tissues, based on engineered tissues and regenerative bioactive materials will become a major clinical impact. Regeneration of tissues implies restoration of structure, restoration of function, restoration of metabolic and biochemical behavior, and restoration of biomechanical performance. Our challenge for the future is to extend these findings to studies in compromised bones with osteopenia and osteoporosis, to apply the findings to larger animals, and especially humans, with aging bones, and to use the findings to design the 3-D architectures required for engineering of tissues [349].\nBased on what we have seen so far, we can see further research and development in materials, technology, engineering and clinical applications to provide better health services coming years, particularly characterized by an ever-aging society. In this section, among many future perspectives associated with medical and dental, as well as industry, Ti biomaterials development and related technologies will be revealed.\n\n5.1. Titanium Industry and New Materials R\u0026D\nThe titanium industry is rebounding with projections that the worldwide mill product shipments by 2008 will be about 56,000 tones per year – up significantly from the 43,000 tones per year of 2002/2003. This is a result of the increase in commercial airplane production along with increased sales to the military, industrial, and consumer markets. As always, cost remains the major barrier to ore titanium use, especially in the industrial and consumer markets [350]. Numerous attempts have been undertaken in the last 60 years to reduce the cost of producing titanium [351–353]. With the advent of high-quality, lower-cost titanium powders, the emphasis in titanium powder metallurgy P/M technology has centered on production of near-net shapes with acceptable levels of mechanical properties. The pre-alloyed, blended elemental, and metal injection molding (MIM) approaches are all looking attractive, and should post significant growth in the next few years [354].\nAmong many developed titanium materials, these are just few to list of newly developed titanium materials; V-free Ti alloys [355,356], Ti-Pd-Co alloy [357], Ti-V-Fe-Al alloy [358], Ti-Cr alloy [359], Ti-Cu-Ni-Sn-M (M: Nb, Ta, Mo) alloy [360], Ti-Cu-Pd [361], Ti-Cu-Si [362], Ti-Zr [363], Ti-Hf [364,365].\nAmorphous materials are non-crystalline solids. For the last three decades, amorphous alloys have attracted great interest because of the results from their new alloy compositions and new atomic configurations [366]. TiAl amorphous alloy provides high strength, linear elastic behavior, and the infinite fatigue life necessary for high device reliability. This alloy was originally developed for material for the digital micromirror device (DMD) chip, and actually it is a TiAl3 phase [367]. Bulk amorphous Ti-based alloys were found to be formed in the diameter range up to 5 mm for the Ti-Ni-Cu-Sn and Ti-Ni-Cu-Si-B systems, which possess a high glass-forming ability [366,368]. There are newly reported Ti45-Ni20-Cu25-Sn5-Zr5 [369] and Ti50-Cu20-Ni24-Si4-B2 [370] that can be amorphatized, too.\n\n5.2. Gradient Functional Material System\nMaterial is composed of multilayer, with each layer having unique characteristics, yet adjacent layers having some similarity is called gradient functional material (GFM). Although such functions can include various properties, it is limited to mechanical, physical, or thermal properties since other properties, such as chemical or electrochemical, are more likely important to the surface layer, and not related to bulky or semi-bulky behavior. For example, if hydroxyapatite is needed to spray-coat onto CpTi, this GFM concept can be applied. Instead of applying HA powder directly onto the CpTi surface, a multilayer of the following sequence: HA/HA + Al2O3/Al2O3 + TiO2/TiO2/CpTi can be prepared to enhance the bonding strength. From the HA side to CpTi side, the mechanical properties (particularly, modulus of elasticity) and thermal properties (such as linear coefficient of thermal expansion) are gradually changing, so that when this HA-coated CpTi is subjected to stressing, interfacial stress between each constituent layer can be minimized, resulting in that the degree of discreteness in the stress field can also be minimized.\nA novel technology for forming a gradient functional titanium-oxide film on a titanium alloy (Ti-6Al-4V) was developed by the reactive DC sputtering vapor deposition method [371]. The method was developed for fabricating denture bases and implants, and the oxygen concentration was changed continuously during sputtering to provide a gradient in the film composition, by which adhesivity to the alloy, surface hardness, and biocompatibility was improved. Denture bases produced by superplastic forming (SPF) are cleaned with an organic solvent, the oxygen concentration is changed continuously during sputtering, and pure titanium is vapor-deposited by the reactive DC sputtering. In the initial stage, intermetallic bonding is achieved by oxygen-free vapor deposition, so the adhesion is excellent and there is no fear of exfoliation. But farther away from the metal surface, the oxygen concentration is raised gradually to form a gradient film. At the surface some titanium oxides are formed. Titanium oxide features excellent biocompatibility, and since it is a hard material, resists damage. The overall film thickness in the experimental was 3 μm, and the Vickers hardness of the surface was 1,500 (200–300 for pure titanium) [371].\nBogdanski et al. [372] fabricated the functionally graded material, which was prepared through powder metallurgical processing with thoroughly mixed powders of the elements. Ten mixtures were prepared ranging from Ni:Ti of 90:10 (by atomic %) through 80:20, to Ni:Ti of 10:90, and pure Ti (0:100). The compaction was done by hot isostatic pressing (HIP) at 1,050 °C and 195 MPa for 5 h. It was reported that, using cells (comprised of osteoblast-like osteosarcoma cells, primary human osteoblasts, and murine fibroblasts), good biocompatibility of Ni-Ti alloys has shown up to about Ni 50% [372].\n\n5.3. Coating\nSurface modifications have been applied to metallic biomaterials in order to improve their wear properties, corrosion resistance, and biocompatibility. Methods of applying calcium phosphate-based materials are being actively investigated with the aim of enhancing osteoinduction on titanium materials [373,374]. This work is necessary because a plasma sprayed calcium phosphate coating has disadvantages, such as the need for a critical thickness to ensure complete coverage of the implant surface [375]. Another approach to enhancing osteoinduction is to promote the formation of hydroxyapatite on titanium in the human body. Calcium ion implantation [376] and a CaTiO3 coating [377] for titanium materials have been examined, and the improvement of biocompatibility with bone was confirmed.\nThe development of a post-operative infection following the implantation, such as a Ti-6Al-4V alloy total joint prosthesis, is a severe complication in many orthopedic surgeries. Preventing these bacterial infections could theoretically be accomplished by administering therapeutic doses of antibiotics as close to the implant site as possible. Mixing antibiotics with PMMA (polymethylmetaacrylate) bone cements has been shown to provide adequate local antibiotic concentrations for extended periods of time [378–380]. Because metallic materials dominate orthopedic bioprosthetic devices, there exists a definite need for developing methods to attach antibiotics to metallic surfaces. Since the naturally forming passive surface oxide layer of Ti-6Al-4V is thought to be responsible for the excellent biocompatibility and corrosion resistance of this alloy, this oxide layer would be a natural choice for facilitating antibiotic attachment and retainment. By carefully controlling the surface chemistry of the oxide and utilizing the pH dependence of surface charge characteristics of the oxide, the attachment of charged antibiotics may be facilitated at suitable pH values. Such a concept has already been successfully tested with macroporous oxides (1–10 μm pores) formed in sulfuric acid solutions [381]. Dunn et al. [378] also studied the microporous (about 1.5 μm) anodized oxides formed on Ti-6Al-4V alloy to facilitate the attachment and sustained release of antibiotics for longer times. The degree of entamicin sulfate attachment and retainment to microporous oxide layers created on the surface of Ti-6Al-4V materials was determined to be a function of the oxide morphology and surface chemistry. Sulfuric (5–10%) anodized samples were observed to retain the electrostatically attached antibiotic for a period of 13 days when washed in saline at a pH of 7.4. It was found that a longer retention of gentamicin by potentiostatically anodized surfaces in phosphoric acid may be attributed to the lower isoelectric point and more negative zeta potential of these surfaces [378]. Similar studies were conducted by Kato et al. [382,383] to evaluate the applicability of the titanium material as a carrier or a substratum. Spongy titanium adsorbed bone morphogenetic protein (BMP) was implanted in muscle pouches in the thighs of mice. It was found that the quantity of new bone induced was somewhat less than that of the control.\nThe adsorption of bovine serum albumin (BSA) on titanium powder has been studied as a function of protein concentration and pH, and in the presence of calcium and phosphate ions. Isotherm data have shown that the adsorption process does not follow the Langmuir model (inflection points). For the pH dependence of adsorption, it was found that (i) the amount adsorbed increased with decreasing pH, indicating that hydration effects are important, and (ii) adsorption increases and decreases in the presence of calcium and phosphate ions, indicating that electrostatic effects are important. The time dependence, isotherm, and desorption data provide indirect evidence of possible conformational changes in the BSA molecule [384]. Hence, protein adsorption is a dynamic event with proteins adsorbing and desorbing as a function of time. McAlarney et al. [385] investigated the role of complement C3 in the competitive adsorption of proteins from diluted human plasma (the Vroman effect) onto TiO2 surfaces. Ti oxide surfaces were made: (1) four anatase surfaces (70 nm, 140 nm, 70 nm aged and solid anatase), (2) three rutile surfaces (70 nm, 140 nm, and solid rutile), and (3) one electropolished Ti. It was found that (i) in both rutile and anatase surfaces, there was an increase in adsorption with increasing oxide film thickness and/or crystallinity, and (ii) anatase surfaces had greater C3 concentration than the equivalent rutile surfaces [385].\nTitanium dental implants are widely used with success, but their rejection is not rare. One of the causes for implant failure may be due to biofilms created by interactions between the implant material and the surrounding tissues and fluids. The study described the selective adsorption of a specific salivary protein to Ti-oxide and the mechanism of adsorption. Klinger et al. [386] treated enamel powder, CpTi powder, as well as Ti powder by Ca, Mg, or K, which were suspended in vitro in human clarified whole saliva, or in various concentrations of purified salivary constituents, at pH 3.0 and 7.0. The powders were then suspended in EDTA solution in order to release proteins that may have adsorbed to their surfaces. It was found that (i) Ti powders adsorbed considerably less salivary proteins as compared with the enamel powder, (ii) human salivary albumin was identified by Western-immunoblot as the main protein that adsorbed to Ca-treated Ti powder, (iii) the Ca effect was not evident at pH 3.0 due to a neutral-basic shift of the protein at a pH level lower than its isoelectric point, and (iv) the in vivo investigation of salivary proteins adsorbing to Ti parts confirmed these results. Based on these findings, it was concluded that albumin was shown to be the main salivary protein adsorbing to Ti via a selective calcium and pH-dependent mechanism, and these findings are important for the understanding of Ti biocompatibility properties, as well as patterns of bacterial dental plaque accumulation on Ti implants, and the consequent implant success [386].\nHayakawa et al. [387] investigated to attach fibronectin directly to a titanium surface treated with tresyl chloride (2,2,2-trifluoroethanesulfonyl chloride) for the development of a strong connection of a dental implant to subepithelial connective tissues and/or peri-implant epithelia. Basic terminal OH groups of mirror polished titanium were allowed to react with tresyl chloride at 37 °C for 2 days. After the reaction of fibronectin with titanium, the X-ray photoelectron spectroscopy revealed the remarkable effect of the activation of terminal OH groups with the tresyl chloride treatment. It was mentioned that fibronectin, a well-known cell-adhesive protein, could easily be attached to the titanium surface by use of the tresyl chloride activation technique [387]. Studies in developmental and cell biology have established the fact that responses of cells are influenced to a large degree by morphology and composition of the extracellular matrix. In order to use this basic principle for improving the biological acceptance of implants by modifying the surfaces with components of the extracellular matrix (ECM), Bierbaum et al. [388,389] modified titanium surfaces with the collagen types I and III in combination with fibronectin. It was reported that (i) increasing the collagen type III amount resulted in a decrease of fibril diameter, while no significant changes in adsorption could be detected, (ii) the amount of fibronectin bound to the heterotypic fibrils depended on fibrillogenesis parameters, such as ionic strength or concentration of phosphate, and varied with the percentage of integrated type III collagen, and (iii) the initial adhesion mechanism of the cells depended on the substrate (titanium, collagen, fibronectin) [388,389].\nCollagen, as a major constituent of human connective tissues, has been regarded as one of the most important biomaterials. Kim et al. [390] investigated the fibrillar self-assembly of collagen by incubating acid-dissolved collagen in an ionic-buffered medium at 37 °C. It was reported that (i) the degree of assembly was varied with the incubation time and monitored by the turbidity change, (ii) the partially assembled collagen contained fibrils with varying diameters, as well as nonfibrillar aggregates, while the fully assembled collagen showed the complete formation of fibrils with uniform diameters of approximately 100–200 nm with periodic stain patterns within the fibrils, which are typical of native collagen fibers, and (iii) without the assembly, the collagen layer on Ti adversely affected the cell attachment and proliferation [390].\nA unique surface treatment on Ti was developed by Wang et al. [391]. Titanium screws and titanium flat sheets were implanted into the epithelial mantle pearl sacs of a fresh water bivalve by replacing the pearls. After 45 days of cultivation, the implant surfaces were deposited with a nacre coating with iridescent luster. The coating could conform, to some extent, to the thread topography of the screw implant, and was about 200–600 μm in thickness. It was found that (i) the coating was composed of a laminated nacreous layer and a transitional non-laminated layer that consisted mainly of vaterite and calcite polymorphs of calcium carbonate, and (ii) the transitional layer was around 2–10 μm thick in the convex and flat region of the implant surface, and could form close contact with titanium surface while the transitional layer was much thicker in the steep concave regions, and could not form close contact with the titanium surface. It was hence concluded that it was possible to fabricate a biologically active and degradable, and mechanically tough and strong nacre coating on titanium dental implants [391].\nFrosch et al. [392] evaluated the partial surface replacement of a knee with stem cell-coated titanium implants for a successful treatment of large osteochondral defects. Mesenchymal stem cells (MSCs) were isolated from bone marrow aspirates of adult sheep. Round titanium implants were seeded with autologous MSC and inserted into an osteochondral defect in the medial femoral condyle. As controls, defects received either an uncoated implant or were left untreated. Nine animals with 18 defects were sacrificed after six months. It was reported that (i) the quality of regenerated cartilage was assessed by in situ hybridization of collagen type II and immunohistochemistry of collagen types I and II, (ii) in 50% of the cases, defects treated with MSC-coated implants showed a complete regeneration of the subchondral bone layer, (iii) a total of 50% of MSC-coated and uncoated implants failed to osseointegrate, and formation of fibrocartilage was observed, (iv) untreated defects, as well as defects treated with uncoated implants, demonstrated incomplete healing of subchondral bone and formation of fibrous cartilage. It was, therefore, concluded that in a significant number of cases, a partial joint resurfacing of the knee with stem cell-coated titanium implants occurs, and a slow bone and cartilage regeneration and an incomplete healing in half of the MSC-coated implants are limitations of the method [392]. The osseointegration of four different kinds of bioactive ceramic-coated Ti screws were compared with uncoated Ti screws by biomechanical and histomorphometric analysis by Lee et al. [152]. Calcium pyrophosphate, 1:3 patite-wollastonite glass ceramic, 1:1 apatite-wollastonite glass ceramic, and bioactive CaO-SiO2-B2O3 glass ceramic coatings were prepared and coated by the dipping method. Coated and uncoated titanium screws were inserted into the tibia of 18 adult mongrel male dogs for 2, 4, and 8 weeks. It was reported that (i) at 2 and 4 weeks after implantation, the extraction torque of ceramic-coated screws was significantly higher than that of uncoated screws, (ii) at 8 weeks, the extraction torques of calcium pyrophosphate coated and both apatite-wollastonite glass ceramics-coated screws were significantly higher than those of CaO-SiO2-B2O3 glass-coated and uncoated screws, and (iii) the fixation strength was increased by the presence of osteoconductive coating materials, such as calcium pyrophosphate, and apatite-wollastonite glass ceramic, which enabled the achievement of higher fixation strength even as early as 2–8 weeks after the insertion [152].\nBigi et al. [393] performed a fast biomimetic deposition of hydroxyapatite (HA) coatings on Ti-6Al-4V substrates using a slightly supersaturated Ca/P solution, with an ionic composition simpler than that of simulated body fluid (SBF) to fabricate nanocrystalline HA. It was found that (i) soaking in supersaturated Ca/P solution results in the deposition of a uniform coating in a few hours, whereas SBF, or even 1.5 × SBF, requires 14 days to deposit a homogeneous coating on the same substrates, (ii) the coating consists of HA globular aggregates, which exhibit a finer lamellar structure than those deposited from SBF, and (iii) the extent of deposition increases on increasing the immersion time [393].\n\n5.4. Fluoride Treatment\nIt is known in the literature that fluoride ions have osteopromoting capacity leading to increased calcification of the bone. Titanium fluoride is reported to form a stable layer on enamel surfaces consisting of titanium which share the oxygen atoms of phosphate on the surface of hydroxyapatite. Ellingsen et al. [394] investigated as to whether a similar, or rather reverse, reaction would take place on fluoride pre-treated titanium after implantation in bone. Threaded TiO2-blasted titanium implants were pre-conditioned with fluoride. The implants were operated into the tibia of Chinchilla rabbits and let to heal for two months before sacrificing the animals. The strength of the bonding between the implants and bones was tested by removing the implants from the bones by the use of an electronic removal torque gauge. It was reported that (i) the fluoride conditioned titanium implants had a significantly increased retention in bone (69.5 N-cm) compared to non-treated blasted implants (56.0 N-cm) and smooth surface implants (17.2 N-cm), and (ii) the histological evaluation revealed that new bones formed on the surface of the test implants, as well as in the marrow or cancellous regions, which was not observed in the control groups, suggesting that fluoride conditioning of titanium has an osteopromoting effect after implantation [394]. Furthermore, push-out tests of fluoridated and control Ti implants placed in rabbits for up to 8 weeks were conducted [395]. It was found that the fluoridated implants sustained greater push-out forces than the controls, and substantial bone adhesion was observed in fluoridated implants, whereas the controls always failed at the interface between the bone and foreign materials. In other rabbit test conducted by Ellingsen et al. [396], it was reported that the fluoridated, blasted implants showed a significantly higher removal torque than the blasted test implant, again indicative of a bioactive reaction of fluoridated Ti implants.\n\n5.5. Laser Applications\nAs mentioned before, laser technology and laser application have advanced remarkably. Lasers can be made to heat, melt, or vaporize materials, depending on laser power density [397,398]. Materials absorb power more readily from Nd:Yag laser beams (λ = 1.06 μm) than they do from CO2 laser beams (λ = 10.6μm). By heating materials, materials can be annealed, or solid state phase-transformation hardened. Using melting, materials can be alloyed, cladded, grain refined, amorphatized, and welded. Using vaporization, materials can be thin film deposited, cleaned, textured, and etched. Using shocking, materials can be shock hardened/peened [399]. Recent advances in the Nd:Yag and CO2 lasers have made possible a wide range of applications and emerging technologies in the computer, microelectronics, and materials fields. In the realm of materials processing, surface treatments and surface processing, surface treatments and surface modification for metals and semiconductors are of particular interest. With appropriate manipulation of the processing conditions (e.g., laser power density or interaction time), a single laser can be used to perform several processes [400,401].\nLaser alloying is a material-processing method that utilizes the high power density available from focused laser sources to melt metal coatings and a portion of the underlying substrate [402]. Since the melting occurs in a very short time and only at the surface, the bulk of the material remains cool, thus serving as an infinite heat sink. Large temperature gradients exist across the boundary between the melted surface region and the underlying solid substrate. This results in rapid self-quenching (1011 ks−1) and resolidification (velocities of 20 m/s). What makes laser surface alloying both attractive and interesting is the wide variety of chemical and microstructural states that can be retained because of the repaid quench from the liquid phase. The types of observed microstructures include extended solid solutions, metastable crystalline phases and metallic glasses as an amorphous metal [402,403]. Alloy production with a wide variety of elements, as well as a wide range of compositional content, can also be accomplished by a mechanical alloying or powder metallurgy, both of which do not involve liquids (in other words, in solid state fabrication).\nDirect laser forming (DLF) is a rapid prototyping technique that enables prompt modeling of metal parts with high bulk density on the base of individual three-dimensional data, including computer tomography models of anatomical structures. Hollander et al. [404] investigated DLFed Ti-6Al-4V for its applicability as hard tissue biomaterial. It was reported that rotating bending tests revealed that the fatigue profile of post-DLF annealed Ti-6Al-4 V was comparable to cast/hot isostatic pressed alloy. In an in vitro investigation, human osteoblasts were cultured on non-porous and porous blasted DLFed Ti-6Al-4V specimens to study morphology, vitality, proliferation and differentiation of the cells. It was reported that (i) the cells spread and proliferated on DLFed Ti-6Al-4V over a culture time of 14 days, (ii) on porous specimens, osteoblasts grew along the rims of the pores and formed circle-shaped structures, as visualized by live/dead staining, as well as scanning electron microscopy, and (iii) overall, the DLFed Ti-6Al-4V approach proved to be efficient, and could be further advanced in the field of hard tissue biomaterials [404].\nRecently, the femtosecond-laser-based tooth preparation technique has been developed [405,406]. Any one of the existing laser technologies using a CO2 laser, Er:Yag laser, Ho:Yag laser, excimer laser, frequency-doubled Alexandrite laser, superpulsed CO2 laser, or picosecond Nd:Yag laser induce severe thermal adverse effects or do not supply sufficient ablation rates for completion of the mechanical drill [407]. Using the femtosecond laser technology for micromachining was successfully developed, for example in machining tools for repairing photolithographic masks or fuel injector nozzles [408].\n\n5.6. Near-Net Shape (NNS) Forming\nIn order to achieve the better condition for fit, for example superstructure for implants or denture base for prostheses, the accuracy of the final products are very crucial and need to be improved. Among various manufacturing technologies, near-net shape forming and nanotechnology are most promising and supportive for these specific aims. New classes of fabrication processes, such as direct-write (DW) technique [409] and solid freedom fabrication (SFF) [410], have established the capability to produce 3-dimensional parts faster, cheaper, and with added functionality. The choice of starting materials and the specific processing technique will produce unique microstructures that impact the final performance, especially of macroscopic structural and electronic parts. Furthermore, the ability to do point-wise deposition of one or more materials provides the opportunity for fabricating structures with novel microstructural and macrostructural features, such as micro-engineered porosity, graded interfaces, and complex multi-material constructions. Near-net shape (NNS) processing offers cost reduction by minimizing machining, reducing part count, and avoiding part distortion from welding. NSS technologies such as flow-forming (FF), superplastic forming (SPF), casting, forging, powder metallurgy (P/M) methods, three-dimensional laser deposition, and plasma arc deposition (PAD) have been explored for potential use in tubular geometries and other shapes of varying complexity. It appears that the reduction in cost of a given titanium product will be maximized by achieving improvements in all of the manufacturing steps, from extraction to finishin [411].\n\n5.7. Tissue Engineering and Scaffold Structure and Materials\nTissue engineering can perhaps be best defined as the use of a combination of cells, engineering materials, and suitable biochemical factors to improve or replace biological functions [412]. The advanced medicine indicates an interdisciplinary field that applies the principles of engineering and life sciences toward the development of biological substitutes that restore, maintain, or improve tissue function for understanding the principles of tissue growth, and applying this to produce functional replacement tissue for clinical use [412]. The term “regenerative medicine” is often used synonymously with “tissue engineering”, although those involved in regenerative medicine place more emphasis on the use of stem cells to produce tissues [412]. Tissue engineering in vitro and in vivo involves the interaction of cells with a material surface. The nature of the surface can directly influence cellular response, ultimately affecting the rate and quality of new tissue formation. Initial events at the surface include the orientated adsorption of molecules from the surrounding fluid, creating a conditioned interface to which the cell responds. The gross morphology, as well as the micro-topography and chemistry of the surface, determine which molecules can adsorb and how cells will attach and align themselves. The local attachments made of the cells with their substrate determine cell shape, which, when transduced via the cytoskeleton to the nucleus, result in expression of specific phenotypes. Osteoblasts and chondrocytes are sensitive to subtle differences in surface roughness and surface chemistry. Boyan et al. [413] investigated the chondrocyte response to TiO2 of differing crystallinities, and showed that cells can discriminate between surfaces at this level as well. Cellular response also depends on the local environmental and state of maturation of the responding cells. It was mentioned that optimizing surface structure for site-specific tissue engineering is one option; modifying surfaces with biological means is another biological engineering [413].\nOne major determination of the suitability of various engineering materials for use in biological settings is the relative strength of adhesion obtained between those materials and their contacting viable phases [414]. Maximal adhesive strength and immobility are desired for orthopedic and dental implants. For example, while minimal bio-adhesion is critical to preventing unwanted thrombus formation in cardiovascular devices, plaque buildup on dental prostheses, and bacterial fouling [414]. Attention should be directed to adhesive phenomena in the oral environment, examining new surface conditioning methods for the prevention of micro-organism deposits, as well as the promotion of excellent tissue bonding to implanted prosthetic devices. Other bio-adhesive phenomena considered included those important to the safe and effective function of new cardiovascular devices [414].\nScaffold material has a two-fold function: artificial extracellular matrices (ECM) and as a spacer keeping a certain open space [415]. Furthermore, scaffold material has to be dissolved completely into the living body after auto-cell is regenerated with artificial extracellular matrices [415]. There are several important biodegradable and/or biofunctional scaffold architectures, structures and materials. They include blended-polymer scaffolds, collagen-based scaffolds, and composite scaffolds of polyhydroxybutyrate-polyhydroxyvalerate with bioactive wollastonite (CaSiO3) [416]. Using an ink-injection technique [417], a thin film (with thickness of about 0.1mm) of calcium phosphate and binding agent is injected onto the substrate to build 3-D bony-like structures [418]. Lee et al. [419] employed three-dimensional printing (3DP) technology to fabricate porous scaffolds by inkjet printing liquid binder droplets. Direct 3DP, where the final scaffold materials are utilized during the actual 3DP process, imposes several limitations on the final scaffold structure. An indirect 3DP protocol was developed, where molds are printed and the final materials are cast into the mold cavity to overcome the limitations of the direct technique. Results of SEM observations revealed highly open, well interconnected, uniform pore architecture (about 100–150 μm) [419]. Scaffold materials for bone tissue engineering often are supplemented with bone morphogenetic proteins (BMPs). Walboomers et al. [420] investigated a bovine extracellular matrix product containing native BMPs. Hollow cylindrical implants were made from titanium fiber mesh, and were implanted subcutaneously into the back of Wistar rats. It was reported that (i) after eight weeks, in two out of six loaded specimens, newly-formed bone and bone marrow-like tissues could be observed, and (ii) after 12 weeks, this had increased to five out of six loaded samples. It was, therefore, concluded that the bovine extracellular matrix product loaded in a titanium fiber mesh tube showed bone-inducing properties [420].\nElectrospinning [421] has recently emerged as a leading technique for generating biomimetic scaffolds made of synthetic and natural polymers for tissue engineering applications. Li et al. [422] compared collagen, gelatin (denatured collagen), solubilized alpha-elastin, and recombinant human tropoelastin as biopolymeric materials for fabricating tissue engineered scaffolds by electrospinning. It was reported that (i) the average diameter of gelatin and collagen fibers could be scaled down to 200–500 nm without any beads, while the alpha-elastin and tropoelastin fibers were several microns in width, and (ii) cell culture studies confirmed that the electrospun engineered protein scaffolds support attachment and growth of human embryonic palatal mesenchymal cells [422]. For fabricating meshes of collagen and/or elastin by means of electrospinning from aqueous solutions, Buttafoco et al. [423] added polyethylene oxide and NaCl to spin continuous and homogeneous fibers. It was reported that (i) upon crosslinking, polyethylene oxide and NaCl were fully leached out, and (ii) smooth muscle cells grew as a confluent layer on top of the crosslinked meshes after 14 days of culture [423].\nSurface properties of scaffolds play an important role in cell adhesion and growth. Biodegradable poly(α-hydroxy acids) have been widely used as scaffolding materials for tissue engineering; however, the lack of functional groups is a limitation. Liu et al. [424] mentioned in their studies that gelatin was successfully immobilized onto the surface of poly(α-hydroxy acids) films and porous scaffolds by an entrapment process. It was found that (i) the amount of entrapped gelatin increased with the ratio of dioxane/water in the solvent mixture used, (ii) chemical crosslinking after physical entrapment considerably increased the amount of retained gelatin on the surface of poly(α-hydroxy acids), (iii) osteoblasts were cultured on these films and scaffolds, (iv) cell numbers on the surface-modified films and scaffolds were significantly higher than those on controls 4 h and 1 day after cell seeding, (v) the osteoblasts showed higher proliferation on surface-modified scaffolds than on the control during 4 weeks of in vitro cultivation, and (vi) more collagen fibers and other cell secretions were deposited on the surface-modified scaffolds than on the control scaffolds [424].\nThere are still unique scaffold systems developed, such as the collagen-carbon nanotubes composite matrices [425], chitosan-based hyaluronan hybrid polymer fibers system [426], bioactive porous CaSiO3 scaffold structure [427], or a three-dimensional porous scaffold composed of biodegradable polyesters [428].\n\n5.8. Application of Nanotechnology to Surface Modification\nSeveral studies have suggested that materials with nanopatterned surfaces produced from various chemistries, such as metals, polymers, composites and ceramics, exhibit better osseointegration when compared to conventional materials [429–432]. Nano-patterned surfaces provide a higher effective surface area and nanocavities when compared to the conventional microrough surfaces. These properties are crucial for the initial protein adsorption that is very important in regulating the cellular interactions on the implant surface. Taylor [433] pointed out that nanotechnology offers the key to faster and remote diagnostic techniques – including new high throughput diagnostics, multi-parameter, tunable diagnostic techniques, and biochips for a variety of assays. It also enables the development of tissue-engineered medical products and artificial organs, such as heart valves, veins and arteries, liver, and skin. These can be grown from the individual’s own tissues as stem cells on a 3-D scaffold, 3-D tissue engineering extracellular matrix, or the expansion of other cell types on a suitable substrate. The applications which seem likely to be most immediately in place are external tissue grafts; dental and bone replacements; protein and gene analysis; internal tissue implants; and nanotechnology applications within in vivo testing devices and various other medical devices. Nanotechnology is applied in a variety of ways across this wide range of products. Artificial organs will demand nanoengineering to affect the chemical functionality presented at a membrane or artificial surface, and thus avoid rejection by the host. There has been much speculation and publicity about more futuristic developments such as nanorobot therapeutics, but these do not seem likely within our time horizon [433].\nThere are studies done on nanotubes. For example, Frosch et al. [434] investigated the effect of different diameters of cylindrical titanium channels on human osteoblasts. Titanium samples with continuous drill channels with various diameters (300, 400, 500, 600, and 1,000 microns) were put into osteoblast cell cultures that were isolated from 12 adult human trauma patients. It was reported that (i) within 20 days, cells grew an average of 838 μm into the drill channels with a diameter of 600 μm, and were significantly faster than in all other channels, (ii) cells produced significantly more osteocalcin messenger RNA (mRNA) in 600 μm channels than they did in 1,000 μm channels, and demonstrated the highest osteogenic differentiation, (iii) the channel diameter did not influence collagen type I production, and (iv) the highest cell density was found in 300 μm channels, suggesting that the diameter of cylindrical titanium channels has a significant effect on migration, gene expression, and mineralization of human osteoblasts [434]. Macak et al. [435] reported on the fabrication of self-organized porous oxide-nanotube layers on the biomedical titanium alloys Ti-6Al-7Nb and Ti-6Al-4V by an anodizing treatment in 1M (NH4)2SO4 electrolytes containing 0.5 wt % of NH4F. It was shown that (i) under specific anodization conditions, self-organized porous oxide structures can be grown on the alloy surface, (ii) SEM images revealed that the porous layers consist of arrays of single nanotubes with a diameter of 100 nm and a spacing of 150 nm, (iii) for the V-containing alloy, enhanced etching of the β-phase is observed, leading to selective dissolution and an inhomogeneous pore formation, and (iv) for the Nb-containing alloy an almost ideal coverage of both phases is obtained. According to XPS measurements, the tubes are a mixed oxide with an almost stoichiometric oxide composition, and can be grown to thicknesses of several hundreds of nanometers, suggesting that a simple surface treatment for Ti alloys has high potential for biomedical applications [435]. A vertically aligned nanotube array of titanium oxide was fabricated on the surface of titanium substrate by anodization. The nanotubes were then treated with NaOH solution to make them bioactive, and to induce growth of hydroxyapatite (bone-like calcium phosphate) in a simulated body fluid. It is found that (i) the presence of TiO2 nanotubes induces the growth of a “nano-inspired nanostructure”, i.e., extremely fine-scale (∼8 nm feature) nanofibers of bioactive sodium titanate structure on the top edge of the ∼15 nm thick nanotube wall, (ii) during the subsequent in vitro immersion in a simulated body fluid, the nano-scale sodium titanate, in turn, induced the nucleation and growth nano-dimensioned hydroxyapatite phase, and (iii) such TiO2 nanotube arrays and associated nanostructures can be useful as a well-adhered bioactive surface layer on Ti implant metals for orthopedic and dental implants, as well as for photocatalysts and other sensor applications [436].\nWebster et al. have suggested that enhanced vitronectin adsorption, conformation and bioactivity are the major reason for increased osteoblast adhesion on nanophase alumina [437]. In the last years, a new method has been described to fabricate nanotubular structures on titanium surfaces. These titania nanotubes can be produced by a variety of methods including electrochemical deposition, sol-gel method, hydrothermal processes and anodic oxidation [438–440]. Using this novel approach, several studies showed that the presence of the nanotube structure on a titanium surface induced a significant increase in the action of osteoblastic cells compared to those grown on flat titanium surfaces [441,442]. Nanotubular TiO2 layer produced using anodization has an amorphous crystal structure and it has been shown that using heat-treatment it can be transformed into anatase to improve cellular interactions [441]. In this study, a sintering protocol at 450 °C for 2 h was used to perform a crystal phase transformation of nanotubes. Significantly higher cell proliferation rates and better cellular morphologies were observed on anatase nanotubular surfaces after 7 days of culture, as shown in the literature [442]. Park et al. [443] produced nanotubular surfaces having pore diameter of 15, 20, 30, 50, 70 and 100 nm without heat treatment and documented that on nanotubular surfaces above 50 nm, the cell attachment and spreading was significantly decreased, thereby causing an increased programmed cell death. They only showed better cell proliferation and matrix mineralization results on nanotubes having 15 nm pore diameter. Whereas, in another study, Oh et al. [444] performed heat treatment following anodization and showed that anatase nanotubes having larger pore diameter (70 to 100 nm) MSCs elongate better and undergo selective differentiation into osteoblast-like cells compared to small nanotubes. In the present study, nanotubes having pore diameter of 70–100 nm were produced and impaired cellular functions on non heat-treated amorphous nanotubular surfaces were observed. However, on anatase nanotubular surfaces, significantly increased cell proliferation values were recorded after 7 days of culture, as shown in the literature [441,442,444]. Residual fluorine within the amorphous nanotubes following anodization might be the factor for this decreased cellular density, as stated in the literature [441]. Therefore, this study was able to show that heat treatment is essential for the production of nanotubular implant surfaces since it provides a more ideal oxide crystal structure for the spreading and proliferation of the cells.\nBeside microtopographical features, surface wettability and surface free energy are also important parameters influencing cell attachment, proliferation and differentiation [445]. Bauer et al. [446] cultured rat mesenchymal stem cells on nanotubular titanium surfaces having different wettability characteristics and found an increased cell attachment on super-hydrophobic surfaces compared with super-hydrophilic ones. In the present study, samples having different wettability profiles were obtained following roughening, anodization and heat treatment procedures. After anodization and heat treatment, water contact angles decreased gradually. However, due to the changes in surface chemistries following treatments, no correlation was found between surface wettability and cellular functions. Further studies are needed to evaluate the effect of hydrophilicity and surface chemistry of titania nanotubes on protein adsorption and cell responses.\nThree types of bioactive polymethylmethacrylate (PMMA)-based bone cement containing nano-sized titania (TiO2) particles were prepared, and their mechanical properties and osteoconductivity are evaluated by Goto et al. [447]. The three types of bioactive bone cement were un-silanized TiO2, 50 wt%, silanized TiO2 50 wt%, and 60 wt% mixed to PMMA. Commercially available PMMA cement was used as a control. The cements were inserted into rat tibiae and allowed to solidify in situ. After 6 and 12 weeks, tibiae were removed for evaluation of osteoconductivity. It was reported that (i) bone cements using silanized TiO2 were directly apposed to bone, while un-silanized TiO2 cement and PMMA control were not, (ii) the osteoconduction of cement with 60 wt% of silanized TiO2 was significantly better than that of the other cements at each time interval, and (iii) the compressive strength of cement with 60 wt% of silanized TiO2 was equivalent to that of PMMA, indicating that cement with 60 wt% of TiO2 was a promising material for use as a bone substitute [447]. Since it is essential for the gap between the hydroxyapatite coated titanium and juxtaposed bone to be filled out with regenerated bone, promoting the functions of bone-forming cells is desired. In order to improve orthopedic implant performance, Sato et al. [448] synthesized nanocrystalline hydroxyapatite (HA) powders to coat titanium through a wet chemical process. The precipitated powders were either sintered at 1,100 °C for 1 h in order to produce microcrystalline size HA, or were treated hydrothermally at 200 °C for 20 h to produce nanocrystalline HA. These powders were then deposited onto titanium by a room temperature process. It was reported that (i) the chemical and physical properties of the original HA powders were retained when coated on titanium by the room temperature process, (ii) osteoblast adhesion increased on the nanocrystalline HA coatings compared to traditionally used plasma-sprayed HA coatings, (iii) greater amounts of calcium deposition by osteoblasts cultured on Y-doped nanocrystalline HA coatings were observed [448]. With a wide variety of applications, nanotechnology has attracted the attention of researchers as well as regulators and industrialists, including nanodrugs and drug delivery, prostheses and implants, and diagnostics and screening technologies. We can take advantages of availability of ultra-fine nanomicrostructures of solid metals, alloys, powder, fibers, or ceramics to fabricate superplastically formed products.\nIt is indispensable here to mention the minimally invasive dentistry (MID) and minimally invasive surgery (MIS). The MIS concept has been created to allow new thinking and a new approach to dentistry where restoration of a tooth becomes the last treatment decision rather than first consideration as at present. It provides a practical approach to caries preventive measures based on the notion of demineralization and remineralization in a micro-phase in order to retain healthy teeth. The medical model of MID is characterized by (1) reduction in cariogenic bacteria, (2) preventive measures, (3) remineralization of early enamel lesions, (4) minimum surgical intervention of cavitated lesions, and (5) repair of defective restorations [449]. At the same time, it is mentioned that MIS has several advantages: (1) since the surgical area is narrower, damage on surrounding soft tissue can be minimized, (2) post-operation pain can be minimized, (3) hospital time can be shortened, and (4) early rehabilitation can be initiated [450]. These MIs in both dentistry and medicine inevitably require precisely manufactured prostheses in micro-scale or even nano-scale. It is anticipated that the MI-based technologies, as well as MI-oriented technologies, will be advanced in the near future.\n\n5.9. Bioengineering and Biomaterial-Integrated Implant System\nWith the aforementioned supportive technologies, surfaces of dental and orthopedic implants have been remarkably advanced. These applications can include not only ordinal implant system but also miniaturized implants, as well as customized implants. Dental implant therapy has been one of the most significant advances in dentistry in the past 25 years. The computer and medical worlds are both working hard to develop smaller and smaller components. Using a precise, controlled, minimally invasive surgical (MIS) technique, the mini dental implants (MDI) are placed into the jawbone. The heads of the implants protrude from the gum tissue and provide a strong, solid foundation for securing the dentures. It is a one-step procedure that involves minimally invasive surgery, no sutures, and none of the typical months of healing. Advantages associated with the MDI are (1) it can provide immediate stabilization of a dental prosthetic appliance after a minimally invasive procedure. (2) It can be used in cases where traditional implants are impractical, or when a different type of anchorage system is needed. (3) Healing time required for mini-implant placement is typically shorter than that associated with conventional 2-stage implant placement and the accompanying aggressive surgical procedure. According the clinical reports, a biometric analysis of 1,029 MDI min-implants, five months to eight years in vivo showed that the MDI mini-implant system can be implemented for long-term prosthesis stabilization, and delivers a consistent level of implant success [451].\nIn addition to the aforementioned miniature implants, an immediate loading, as well as customized implants, have been receiving attention recently. Conventionally, a dental implant patient is required to have two-stages of treatment consisting of two dental appointments five to six months apart. Recently, a single stage treatment has received attention. Placing an implant immediately or shortly after tooth extraction offers several advantages for the patient as well as for the clinician. These advantages include shorter treatment time, less bone sorption, fewer surgical sessions, easier definition of the implant position, and better opportunities for osseointegration because of the healing potential of the fresh extraction site [452–455]. Titanium bar (particularly the portions in direct contact to connective tissue and bony tissue) is machined to have the exact shape of the root portion of the extracted tooth of the patient. The expected outcome of this method is a perfect mechanical retention, and therefore an ideal osseointegration can be achieved. This is called a custom (or customized) implant, which is fabricated by the electro-discharge machining (EDM) technique.\nPrefabricated dental implant can be further machined to replicate the extracted tooth and machined implant (whose shape is exactly same as the patient’s extracted tooth) can be prepared within a relatively short operation time and can be placed within one hour to the patient. Since the root shape of the placed implant is exactly same as the extracted tooth’s root form, the follow-up reaction including osseointegration is expected shorter and placed implant’s retention force can be achieved within a relatively short time.\nElectron-beam machining (EBM) is a machining process where high-velocity electrons (in the range of 50 to 200 kV to accelerate electrons to 200,000 km/s) are directed toward a work piece, creating heat and vaporizing the material. Electromagnetic lenses are used to direct the electron beam, by means of deflection, into a vacuum. The electrons strike the top layer of the work piece, removing material, and then become trapped in some layer beneath the surface. Applications of this process are annealing, metal removal, and welding. EBM can be used for very accurate cutting of a wide variety of metals. Surface finish is better and kerf width is narrower than those for other thermal cutting processes. The process is similar to laser-beam machining, but because EBM requires a vacuum, it is not used as frequently as laser-beam machining [456]. In addition to the immediate placement of dental implants, another concept has been introduced. Techniques such as stereoscopic lithography and computer-assisted design and manufacture (CAD/CAM) have been successfully used with computer-numerized control milling to manufacture customized titanium implants for single-stage reconstruction of the maxilla, hemimandible, and dentition without the use of composite flap over after the removal of tumors [457]. Nishimura et al. [458] applied this concept to dental implants to fabricate the individual and splinted customized abutments for all restoration of implants in partially edentulous patients. It was claimed that complicated clinical problems such as angulation, alignment, and position can be solved. However, with this technique, the peri-implant soft tissues are allowed to heal 2 to 3 weeks, so that at least two dental appointments are required.\nMany oral implant companies (about 25 companies are currently marketing 100 different dental implant systems) have recently launched new products with claimed unique, and sometimes bioactive surfaces [459,460]. The focus has shifted from surface roughness to surface chemistry and a combination of chemical manipulations on the porous structure. To properly explain the claims for new surfaces, it is essential to summarize current opinions on bone anchorage, with emphasis on the potentials for biochemical bonding. There were two ways of implant anchorage or retention: mechanical and bioactive [459,460].\nRecent research has further redefined the retention means of dental implants into the terminology of osseointegration versus biointegration. When examining the interface at a higher magnification level, Sundgren et al. [30] showed that unimplanted Ti surfaces have a surface oxide (TiO2) with a thickness of about 35 nm. During an implantation period of eight years, the thickness of this layer was reported to increase by a factor of 10. Furthermore, calcium, phosphorous, and carbon were identified as components of the oxide layer, with the phosphorous strongly bound to oxygen, indicating the presence of phosphorous groups in the metal oxide layer. Many retrospective studies on retrieved implants, as well as clinical reports, confirm the aforementioned important evidence (1) surface titanium oxide film grows during the implantation period, and (2) calcium, phosphorous, carbon, hydroxyl ions, proteins, etc. are incorporated in an ever-growing surface oxide even inside the human biological environments [460,461]. Numerous in vitro studies, (e.g., [461]) on treated or untreated titanium surfaces were covered and to some extent were incorporated with Ca and P ions when such surfaces were immersed in SBF (simulated body fluid). Additionally, we know that bone and blood cells are rugophilia, hence in order not only to accommodate for the bone growth, but also to facilitate such cells adhesion and spreading, titanium surfaces need to be textured to accomplish and show appropriate roughness [462]. Furthermore, gradient functional concept (GFC) on materials and structures has been receiving special attention not only in industrial applications, but in dental as well as medical fields [462]. Particularly, when such structures and concepts are about to be applied to implants, its importance becomes more clinically crucial. For example, the majority of implant mass should be strong and tough, so that occlusal force can be smoothly transferred from the placed implant to the receiving hard tissue [462]. However, the surface needs to be engineered to exhibit some extent of roughness. From such macro-structural changes from bulk core to the porous case, again the structural integrity should be maintained. The GFC can also be applied for the purpose of having a chemical (compositional) gradient. Ca-, P-enrichment is not needed in the interior materials of the implants. Some other modifications related to chemical dressing or conditioning can also be utilized for achieving gradient functionality on chemical alternations on surfaces as well as near-surface zones [462].\n\n5.10. Technology-Integrated Implant Systems\nAny new type of implant (not only dental but also orthopedic applications) should possess a gradual function of mechanical and biological behaviors, so that mechanical compatibility and biological compatibility can be realized with s single implant system [462]. Since microtextured Ti surfaces [67,463,464] and/or porous Ti surfaces [465–467] promote fibroblast apposition and bone ingrowth, the extreme left side representing the solid Ti implant body should have gradually increased internal porosities toward to the case side which is in contact with vital hard/soft tissue. Accordingly, mechanical strength of this implant system decreases gradually from core to case, whereas biological activity increases from core to case side. Therefore, the mechanical compatibility can be completely achieved. Porosity-controlled surface zones can be fabricated by an electrochemical technique [468], polymeric sponge replication method [65], powder metallurgy technique, superplastic diffusion bonding method [469], or foamed metal structure technique [470].\nOnce the Ti implant is placed in hard tissue, TiO2 grows and increases its thickness [30,41,471–478], due to more oxygen availability inside the body fluid, as well as co-existence of superoxidant. It is very important to mention here that Ti is not in contact with the biological environment, but rather there is a gradual transition from the bulk Ti material, stoichiometric oxide (i.e., TiO2), hydrated polarized oxide, adsorbed lipoproteins and glycolipids, portroglycnas, collagen filaments and bundles to cells [40]. Such gradient functional structure can be also fabricated in CpTi and microtextured polyethylene terephthalate (PET) system [479]. In addition, a gradient structural system of Ti and TiN was developed [480]. During HA coating, a gradient functional layer was successfully fabricated [214]. To promote these gradient functional (GF) and gradient structural (GS) transitions, there are many in vivo, as well as in vitro, evidences indicting that surface titanium oxide is incorporated with mineral ions, water and other constituents of biofluids [30,38,39,481]. Since a surface layer of TiO2 is negatively charged, the calcium ion attachment can be easily achieved [33,482]. Retrieved Ti implants showed that surface TiO2 was incorporated with Ca and P ions [483], while in vitro treatment of TiO2 in extracellular fluids or simulated body fluid (SBF) for prolonged periods of incubation time resulted in the incorporation of Ca, P, and S ions into TiO2 [30,38–41,218,471,481,484]. Without prolonged treatment, there are several methods proposed to relatively short-time incubation for incorporation of Ca and P ions. For example, TiO2 can be electrochemically treated in an electrolyte of a mixture of calcium acetate monohydrate and calcium glycerophosphate [145]. As a result of incorporation of Ca and P ions, bone-like hydroxyapatite can be formed in macro-scale [485] or nano-dimension [448]. Again for reducing the incubation time, bone-like hydroxyapatite crystals can be formed by treating the TiO2 surface with water and hydrogen plasma immersion ion implantation, followed by immersion in SBF [146], or by treating in hydrogen peroxide followed by SBF immersion [147], or immersion in SBF while treating the TiO2 surface with micro-arc oxidation and irradiation with UV light [171]. It is also known that P ions can be incorporated into TiO2 while it is immersed in the human serum [40].\nBony growth extreme surface zones should have a same roughness as the roughness of receiving hard tissue through micro-porous texturing techniques. This area can be structured using nanotube concepts [434–436]. Because this zone responds strongly to osseointegration, the structure, as well as the chemistry, should accommodate favorable osteoinductive reactions. Bone morphogenetic protein [385,486,487], and nano-apatite can be coated [488]. The zone may be treated by femtosecond laser machining [404] to build a micro-scale 3D scaffold which is structured inside the macro-porosities. Such scaffold can be made of biodegradable material (e.g., chitosan), which is incorporated with protein, Ca, P, apatite particles or other species possessing bone growth factors [483].\n"}

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Future Perspectives\nResearch and development in science and technology should not be discrete. If we evaluate correctly and examine carefully what have been done in the past, we will be able to foresee what would be available for us in future. Titanium materials science is a typical interdisciplinary science and engineering. In order for engineered titanium materials to serve as titanium biomaterials, we have been discussing and reviewing numerous articles to prove that appropriate surface modifications and characterizations should be properly preformed and reflected to appropriate fabrication technologies and methods.\nIn the past, when tissues became diseased or damaged, a physician had little recourse but to remove the offending part, with obvious limitations. Removal of joints, vertebrae, teeth, or organs led to only a marginally improved quality of life. However, human survivability seldom exceeded the progressive decrease in quality of tissues, so the need for replacement parts was small. During the last century the situation changed greatly. The discovery of antiseptics, penicillin and other antibiotics, chemical treatment of water supplies, improved hygiene, and vaccination all contributed to a major increase in human survivability in developed countries. Life expectancy is now in the range of 80+ years. In the past, it was the major practice to remove the diseased tissues, whereas at the present, either transplants (using autografts, heterografts, or homografts) or implants (using biological fixation, bioactive fixation, or cement fixation) are commonly utilized, and in the future, regeneration of tissues, based on engineered tissues and regenerative bioactive materials will become a major clinical impact. Regeneration of tissues implies restoration of structure, restoration of function, restoration of metabolic and biochemical behavior, and restoration of biomechanical performance. Our challenge for the future is to extend these findings to studies in compromised bones with osteopenia and osteoporosis, to apply the findings to larger animals, and especially humans, with aging bones, and to use the findings to design the 3-D architectures required for engineering of tissues [349].\nBased on what we have seen so far, we can see further research and development in materials, technology, engineering and clinical applications to provide better health services coming years, particularly characterized by an ever-aging society. In this section, among many future perspectives associated with medical and dental, as well as industry, Ti biomaterials development and related technologies will be revealed.\n\n5.1. Titanium Industry and New Materials R\u0026D\nThe titanium industry is rebounding with projections that the worldwide mill product shipments by 2008 will be about 56,000 tones per year – up significantly from the 43,000 tones per year of 2002/2003. This is a result of the increase in commercial airplane production along with increased sales to the military, industrial, and consumer markets. As always, cost remains the major barrier to ore titanium use, especially in the industrial and consumer markets [350]. Numerous attempts have been undertaken in the last 60 years to reduce the cost of producing titanium [351–353]. With the advent of high-quality, lower-cost titanium powders, the emphasis in titanium powder metallurgy P/M technology has centered on production of near-net shapes with acceptable levels of mechanical properties. The pre-alloyed, blended elemental, and metal injection molding (MIM) approaches are all looking attractive, and should post significant growth in the next few years [354].\nAmong many developed titanium materials, these are just few to list of newly developed titanium materials; V-free Ti alloys [355,356], Ti-Pd-Co alloy [357], Ti-V-Fe-Al alloy [358], Ti-Cr alloy [359], Ti-Cu-Ni-Sn-M (M: Nb, Ta, Mo) alloy [360], Ti-Cu-Pd [361], Ti-Cu-Si [362], Ti-Zr [363], Ti-Hf [364,365].\nAmorphous materials are non-crystalline solids. For the last three decades, amorphous alloys have attracted great interest because of the results from their new alloy compositions and new atomic configurations [366]. TiAl amorphous alloy provides high strength, linear elastic behavior, and the infinite fatigue life necessary for high device reliability. This alloy was originally developed for material for the digital micromirror device (DMD) chip, and actually it is a TiAl3 phase [367]. Bulk amorphous Ti-based alloys were found to be formed in the diameter range up to 5 mm for the Ti-Ni-Cu-Sn and Ti-Ni-Cu-Si-B systems, which possess a high glass-forming ability [366,368]. There are newly reported Ti45-Ni20-Cu25-Sn5-Zr5 [369] and Ti50-Cu20-Ni24-Si4-B2 [370] that can be amorphatized, too.\n\n5.2. Gradient Functional Material System\nMaterial is composed of multilayer, with each layer having unique characteristics, yet adjacent layers having some similarity is called gradient functional material (GFM). Although such functions can include various properties, it is limited to mechanical, physical, or thermal properties since other properties, such as chemical or electrochemical, are more likely important to the surface layer, and not related to bulky or semi-bulky behavior. For example, if hydroxyapatite is needed to spray-coat onto CpTi, this GFM concept can be applied. Instead of applying HA powder directly onto the CpTi surface, a multilayer of the following sequence: HA/HA + Al2O3/Al2O3 + TiO2/TiO2/CpTi can be prepared to enhance the bonding strength. From the HA side to CpTi side, the mechanical properties (particularly, modulus of elasticity) and thermal properties (such as linear coefficient of thermal expansion) are gradually changing, so that when this HA-coated CpTi is subjected to stressing, interfacial stress between each constituent layer can be minimized, resulting in that the degree of discreteness in the stress field can also be minimized.\nA novel technology for forming a gradient functional titanium-oxide film on a titanium alloy (Ti-6Al-4V) was developed by the reactive DC sputtering vapor deposition method [371]. The method was developed for fabricating denture bases and implants, and the oxygen concentration was changed continuously during sputtering to provide a gradient in the film composition, by which adhesivity to the alloy, surface hardness, and biocompatibility was improved. Denture bases produced by superplastic forming (SPF) are cleaned with an organic solvent, the oxygen concentration is changed continuously during sputtering, and pure titanium is vapor-deposited by the reactive DC sputtering. In the initial stage, intermetallic bonding is achieved by oxygen-free vapor deposition, so the adhesion is excellent and there is no fear of exfoliation. But farther away from the metal surface, the oxygen concentration is raised gradually to form a gradient film. At the surface some titanium oxides are formed. Titanium oxide features excellent biocompatibility, and since it is a hard material, resists damage. The overall film thickness in the experimental was 3 μm, and the Vickers hardness of the surface was 1,500 (200–300 for pure titanium) [371].\nBogdanski et al. [372] fabricated the functionally graded material, which was prepared through powder metallurgical processing with thoroughly mixed powders of the elements. Ten mixtures were prepared ranging from Ni:Ti of 90:10 (by atomic %) through 80:20, to Ni:Ti of 10:90, and pure Ti (0:100). The compaction was done by hot isostatic pressing (HIP) at 1,050 °C and 195 MPa for 5 h. It was reported that, using cells (comprised of osteoblast-like osteosarcoma cells, primary human osteoblasts, and murine fibroblasts), good biocompatibility of Ni-Ti alloys has shown up to about Ni 50% [372].\n\n5.3. Coating\nSurface modifications have been applied to metallic biomaterials in order to improve their wear properties, corrosion resistance, and biocompatibility. Methods of applying calcium phosphate-based materials are being actively investigated with the aim of enhancing osteoinduction on titanium materials [373,374]. This work is necessary because a plasma sprayed calcium phosphate coating has disadvantages, such as the need for a critical thickness to ensure complete coverage of the implant surface [375]. Another approach to enhancing osteoinduction is to promote the formation of hydroxyapatite on titanium in the human body. Calcium ion implantation [376] and a CaTiO3 coating [377] for titanium materials have been examined, and the improvement of biocompatibility with bone was confirmed.\nThe development of a post-operative infection following the implantation, such as a Ti-6Al-4V alloy total joint prosthesis, is a severe complication in many orthopedic surgeries. Preventing these bacterial infections could theoretically be accomplished by administering therapeutic doses of antibiotics as close to the implant site as possible. Mixing antibiotics with PMMA (polymethylmetaacrylate) bone cements has been shown to provide adequate local antibiotic concentrations for extended periods of time [378–380]. Because metallic materials dominate orthopedic bioprosthetic devices, there exists a definite need for developing methods to attach antibiotics to metallic surfaces. Since the naturally forming passive surface oxide layer of Ti-6Al-4V is thought to be responsible for the excellent biocompatibility and corrosion resistance of this alloy, this oxide layer would be a natural choice for facilitating antibiotic attachment and retainment. By carefully controlling the surface chemistry of the oxide and utilizing the pH dependence of surface charge characteristics of the oxide, the attachment of charged antibiotics may be facilitated at suitable pH values. Such a concept has already been successfully tested with macroporous oxides (1–10 μm pores) formed in sulfuric acid solutions [381]. Dunn et al. [378] also studied the microporous (about 1.5 μm) anodized oxides formed on Ti-6Al-4V alloy to facilitate the attachment and sustained release of antibiotics for longer times. The degree of entamicin sulfate attachment and retainment to microporous oxide layers created on the surface of Ti-6Al-4V materials was determined to be a function of the oxide morphology and surface chemistry. Sulfuric (5–10%) anodized samples were observed to retain the electrostatically attached antibiotic for a period of 13 days when washed in saline at a pH of 7.4. It was found that a longer retention of gentamicin by potentiostatically anodized surfaces in phosphoric acid may be attributed to the lower isoelectric point and more negative zeta potential of these surfaces [378]. Similar studies were conducted by Kato et al. [382,383] to evaluate the applicability of the titanium material as a carrier or a substratum. Spongy titanium adsorbed bone morphogenetic protein (BMP) was implanted in muscle pouches in the thighs of mice. It was found that the quantity of new bone induced was somewhat less than that of the control.\nThe adsorption of bovine serum albumin (BSA) on titanium powder has been studied as a function of protein concentration and pH, and in the presence of calcium and phosphate ions. Isotherm data have shown that the adsorption process does not follow the Langmuir model (inflection points). For the pH dependence of adsorption, it was found that (i) the amount adsorbed increased with decreasing pH, indicating that hydration effects are important, and (ii) adsorption increases and decreases in the presence of calcium and phosphate ions, indicating that electrostatic effects are important. The time dependence, isotherm, and desorption data provide indirect evidence of possible conformational changes in the BSA molecule [384]. Hence, protein adsorption is a dynamic event with proteins adsorbing and desorbing as a function of time. McAlarney et al. [385] investigated the role of complement C3 in the competitive adsorption of proteins from diluted human plasma (the Vroman effect) onto TiO2 surfaces. Ti oxide surfaces were made: (1) four anatase surfaces (70 nm, 140 nm, 70 nm aged and solid anatase), (2) three rutile surfaces (70 nm, 140 nm, and solid rutile), and (3) one electropolished Ti. It was found that (i) in both rutile and anatase surfaces, there was an increase in adsorption with increasing oxide film thickness and/or crystallinity, and (ii) anatase surfaces had greater C3 concentration than the equivalent rutile surfaces [385].\nTitanium dental implants are widely used with success, but their rejection is not rare. One of the causes for implant failure may be due to biofilms created by interactions between the implant material and the surrounding tissues and fluids. The study described the selective adsorption of a specific salivary protein to Ti-oxide and the mechanism of adsorption. Klinger et al. [386] treated enamel powder, CpTi powder, as well as Ti powder by Ca, Mg, or K, which were suspended in vitro in human clarified whole saliva, or in various concentrations of purified salivary constituents, at pH 3.0 and 7.0. The powders were then suspended in EDTA solution in order to release proteins that may have adsorbed to their surfaces. It was found that (i) Ti powders adsorbed considerably less salivary proteins as compared with the enamel powder, (ii) human salivary albumin was identified by Western-immunoblot as the main protein that adsorbed to Ca-treated Ti powder, (iii) the Ca effect was not evident at pH 3.0 due to a neutral-basic shift of the protein at a pH level lower than its isoelectric point, and (iv) the in vivo investigation of salivary proteins adsorbing to Ti parts confirmed these results. Based on these findings, it was concluded that albumin was shown to be the main salivary protein adsorbing to Ti via a selective calcium and pH-dependent mechanism, and these findings are important for the understanding of Ti biocompatibility properties, as well as patterns of bacterial dental plaque accumulation on Ti implants, and the consequent implant success [386].\nHayakawa et al. [387] investigated to attach fibronectin directly to a titanium surface treated with tresyl chloride (2,2,2-trifluoroethanesulfonyl chloride) for the development of a strong connection of a dental implant to subepithelial connective tissues and/or peri-implant epithelia. Basic terminal OH groups of mirror polished titanium were allowed to react with tresyl chloride at 37 °C for 2 days. After the reaction of fibronectin with titanium, the X-ray photoelectron spectroscopy revealed the remarkable effect of the activation of terminal OH groups with the tresyl chloride treatment. It was mentioned that fibronectin, a well-known cell-adhesive protein, could easily be attached to the titanium surface by use of the tresyl chloride activation technique [387]. Studies in developmental and cell biology have established the fact that responses of cells are influenced to a large degree by morphology and composition of the extracellular matrix. In order to use this basic principle for improving the biological acceptance of implants by modifying the surfaces with components of the extracellular matrix (ECM), Bierbaum et al. [388,389] modified titanium surfaces with the collagen types I and III in combination with fibronectin. It was reported that (i) increasing the collagen type III amount resulted in a decrease of fibril diameter, while no significant changes in adsorption could be detected, (ii) the amount of fibronectin bound to the heterotypic fibrils depended on fibrillogenesis parameters, such as ionic strength or concentration of phosphate, and varied with the percentage of integrated type III collagen, and (iii) the initial adhesion mechanism of the cells depended on the substrate (titanium, collagen, fibronectin) [388,389].\nCollagen, as a major constituent of human connective tissues, has been regarded as one of the most important biomaterials. Kim et al. [390] investigated the fibrillar self-assembly of collagen by incubating acid-dissolved collagen in an ionic-buffered medium at 37 °C. It was reported that (i) the degree of assembly was varied with the incubation time and monitored by the turbidity change, (ii) the partially assembled collagen contained fibrils with varying diameters, as well as nonfibrillar aggregates, while the fully assembled collagen showed the complete formation of fibrils with uniform diameters of approximately 100–200 nm with periodic stain patterns within the fibrils, which are typical of native collagen fibers, and (iii) without the assembly, the collagen layer on Ti adversely affected the cell attachment and proliferation [390].\nA unique surface treatment on Ti was developed by Wang et al. [391]. Titanium screws and titanium flat sheets were implanted into the epithelial mantle pearl sacs of a fresh water bivalve by replacing the pearls. After 45 days of cultivation, the implant surfaces were deposited with a nacre coating with iridescent luster. The coating could conform, to some extent, to the thread topography of the screw implant, and was about 200–600 μm in thickness. It was found that (i) the coating was composed of a laminated nacreous layer and a transitional non-laminated layer that consisted mainly of vaterite and calcite polymorphs of calcium carbonate, and (ii) the transitional layer was around 2–10 μm thick in the convex and flat region of the implant surface, and could form close contact with titanium surface while the transitional layer was much thicker in the steep concave regions, and could not form close contact with the titanium surface. It was hence concluded that it was possible to fabricate a biologically active and degradable, and mechanically tough and strong nacre coating on titanium dental implants [391].\nFrosch et al. [392] evaluated the partial surface replacement of a knee with stem cell-coated titanium implants for a successful treatment of large osteochondral defects. Mesenchymal stem cells (MSCs) were isolated from bone marrow aspirates of adult sheep. Round titanium implants were seeded with autologous MSC and inserted into an osteochondral defect in the medial femoral condyle. As controls, defects received either an uncoated implant or were left untreated. Nine animals with 18 defects were sacrificed after six months. It was reported that (i) the quality of regenerated cartilage was assessed by in situ hybridization of collagen type II and immunohistochemistry of collagen types I and II, (ii) in 50% of the cases, defects treated with MSC-coated implants showed a complete regeneration of the subchondral bone layer, (iii) a total of 50% of MSC-coated and uncoated implants failed to osseointegrate, and formation of fibrocartilage was observed, (iv) untreated defects, as well as defects treated with uncoated implants, demonstrated incomplete healing of subchondral bone and formation of fibrous cartilage. It was, therefore, concluded that in a significant number of cases, a partial joint resurfacing of the knee with stem cell-coated titanium implants occurs, and a slow bone and cartilage regeneration and an incomplete healing in half of the MSC-coated implants are limitations of the method [392]. The osseointegration of four different kinds of bioactive ceramic-coated Ti screws were compared with uncoated Ti screws by biomechanical and histomorphometric analysis by Lee et al. [152]. Calcium pyrophosphate, 1:3 patite-wollastonite glass ceramic, 1:1 apatite-wollastonite glass ceramic, and bioactive CaO-SiO2-B2O3 glass ceramic coatings were prepared and coated by the dipping method. Coated and uncoated titanium screws were inserted into the tibia of 18 adult mongrel male dogs for 2, 4, and 8 weeks. It was reported that (i) at 2 and 4 weeks after implantation, the extraction torque of ceramic-coated screws was significantly higher than that of uncoated screws, (ii) at 8 weeks, the extraction torques of calcium pyrophosphate coated and both apatite-wollastonite glass ceramics-coated screws were significantly higher than those of CaO-SiO2-B2O3 glass-coated and uncoated screws, and (iii) the fixation strength was increased by the presence of osteoconductive coating materials, such as calcium pyrophosphate, and apatite-wollastonite glass ceramic, which enabled the achievement of higher fixation strength even as early as 2–8 weeks after the insertion [152].\nBigi et al. [393] performed a fast biomimetic deposition of hydroxyapatite (HA) coatings on Ti-6Al-4V substrates using a slightly supersaturated Ca/P solution, with an ionic composition simpler than that of simulated body fluid (SBF) to fabricate nanocrystalline HA. It was found that (i) soaking in supersaturated Ca/P solution results in the deposition of a uniform coating in a few hours, whereas SBF, or even 1.5 × SBF, requires 14 days to deposit a homogeneous coating on the same substrates, (ii) the coating consists of HA globular aggregates, which exhibit a finer lamellar structure than those deposited from SBF, and (iii) the extent of deposition increases on increasing the immersion time [393].\n\n5.4. Fluoride Treatment\nIt is known in the literature that fluoride ions have osteopromoting capacity leading to increased calcification of the bone. Titanium fluoride is reported to form a stable layer on enamel surfaces consisting of titanium which share the oxygen atoms of phosphate on the surface of hydroxyapatite. Ellingsen et al. [394] investigated as to whether a similar, or rather reverse, reaction would take place on fluoride pre-treated titanium after implantation in bone. Threaded TiO2-blasted titanium implants were pre-conditioned with fluoride. The implants were operated into the tibia of Chinchilla rabbits and let to heal for two months before sacrificing the animals. The strength of the bonding between the implants and bones was tested by removing the implants from the bones by the use of an electronic removal torque gauge. It was reported that (i) the fluoride conditioned titanium implants had a significantly increased retention in bone (69.5 N-cm) compared to non-treated blasted implants (56.0 N-cm) and smooth surface implants (17.2 N-cm), and (ii) the histological evaluation revealed that new bones formed on the surface of the test implants, as well as in the marrow or cancellous regions, which was not observed in the control groups, suggesting that fluoride conditioning of titanium has an osteopromoting effect after implantation [394]. Furthermore, push-out tests of fluoridated and control Ti implants placed in rabbits for up to 8 weeks were conducted [395]. It was found that the fluoridated implants sustained greater push-out forces than the controls, and substantial bone adhesion was observed in fluoridated implants, whereas the controls always failed at the interface between the bone and foreign materials. In other rabbit test conducted by Ellingsen et al. [396], it was reported that the fluoridated, blasted implants showed a significantly higher removal torque than the blasted test implant, again indicative of a bioactive reaction of fluoridated Ti implants.\n\n5.5. Laser Applications\nAs mentioned before, laser technology and laser application have advanced remarkably. Lasers can be made to heat, melt, or vaporize materials, depending on laser power density [397,398]. Materials absorb power more readily from Nd:Yag laser beams (λ = 1.06 μm) than they do from CO2 laser beams (λ = 10.6μm). By heating materials, materials can be annealed, or solid state phase-transformation hardened. Using melting, materials can be alloyed, cladded, grain refined, amorphatized, and welded. Using vaporization, materials can be thin film deposited, cleaned, textured, and etched. Using shocking, materials can be shock hardened/peened [399]. Recent advances in the Nd:Yag and CO2 lasers have made possible a wide range of applications and emerging technologies in the computer, microelectronics, and materials fields. In the realm of materials processing, surface treatments and surface processing, surface treatments and surface modification for metals and semiconductors are of particular interest. With appropriate manipulation of the processing conditions (e.g., laser power density or interaction time), a single laser can be used to perform several processes [400,401].\nLaser alloying is a material-processing method that utilizes the high power density available from focused laser sources to melt metal coatings and a portion of the underlying substrate [402]. Since the melting occurs in a very short time and only at the surface, the bulk of the material remains cool, thus serving as an infinite heat sink. Large temperature gradients exist across the boundary between the melted surface region and the underlying solid substrate. This results in rapid self-quenching (1011 ks−1) and resolidification (velocities of 20 m/s). What makes laser surface alloying both attractive and interesting is the wide variety of chemical and microstructural states that can be retained because of the repaid quench from the liquid phase. The types of observed microstructures include extended solid solutions, metastable crystalline phases and metallic glasses as an amorphous metal [402,403]. Alloy production with a wide variety of elements, as well as a wide range of compositional content, can also be accomplished by a mechanical alloying or powder metallurgy, both of which do not involve liquids (in other words, in solid state fabrication).\nDirect laser forming (DLF) is a rapid prototyping technique that enables prompt modeling of metal parts with high bulk density on the base of individual three-dimensional data, including computer tomography models of anatomical structures. Hollander et al. [404] investigated DLFed Ti-6Al-4V for its applicability as hard tissue biomaterial. It was reported that rotating bending tests revealed that the fatigue profile of post-DLF annealed Ti-6Al-4 V was comparable to cast/hot isostatic pressed alloy. In an in vitro investigation, human osteoblasts were cultured on non-porous and porous blasted DLFed Ti-6Al-4V specimens to study morphology, vitality, proliferation and differentiation of the cells. It was reported that (i) the cells spread and proliferated on DLFed Ti-6Al-4V over a culture time of 14 days, (ii) on porous specimens, osteoblasts grew along the rims of the pores and formed circle-shaped structures, as visualized by live/dead staining, as well as scanning electron microscopy, and (iii) overall, the DLFed Ti-6Al-4V approach proved to be efficient, and could be further advanced in the field of hard tissue biomaterials [404].\nRecently, the femtosecond-laser-based tooth preparation technique has been developed [405,406]. Any one of the existing laser technologies using a CO2 laser, Er:Yag laser, Ho:Yag laser, excimer laser, frequency-doubled Alexandrite laser, superpulsed CO2 laser, or picosecond Nd:Yag laser induce severe thermal adverse effects or do not supply sufficient ablation rates for completion of the mechanical drill [407]. Using the femtosecond laser technology for micromachining was successfully developed, for example in machining tools for repairing photolithographic masks or fuel injector nozzles [408].\n\n5.6. Near-Net Shape (NNS) Forming\nIn order to achieve the better condition for fit, for example superstructure for implants or denture base for prostheses, the accuracy of the final products are very crucial and need to be improved. Among various manufacturing technologies, near-net shape forming and nanotechnology are most promising and supportive for these specific aims. New classes of fabrication processes, such as direct-write (DW) technique [409] and solid freedom fabrication (SFF) [410], have established the capability to produce 3-dimensional parts faster, cheaper, and with added functionality. The choice of starting materials and the specific processing technique will produce unique microstructures that impact the final performance, especially of macroscopic structural and electronic parts. Furthermore, the ability to do point-wise deposition of one or more materials provides the opportunity for fabricating structures with novel microstructural and macrostructural features, such as micro-engineered porosity, graded interfaces, and complex multi-material constructions. Near-net shape (NNS) processing offers cost reduction by minimizing machining, reducing part count, and avoiding part distortion from welding. NSS technologies such as flow-forming (FF), superplastic forming (SPF), casting, forging, powder metallurgy (P/M) methods, three-dimensional laser deposition, and plasma arc deposition (PAD) have been explored for potential use in tubular geometries and other shapes of varying complexity. It appears that the reduction in cost of a given titanium product will be maximized by achieving improvements in all of the manufacturing steps, from extraction to finishin [411].\n\n5.7. Tissue Engineering and Scaffold Structure and Materials\nTissue engineering can perhaps be best defined as the use of a combination of cells, engineering materials, and suitable biochemical factors to improve or replace biological functions [412]. The advanced medicine indicates an interdisciplinary field that applies the principles of engineering and life sciences toward the development of biological substitutes that restore, maintain, or improve tissue function for understanding the principles of tissue growth, and applying this to produce functional replacement tissue for clinical use [412]. The term “regenerative medicine” is often used synonymously with “tissue engineering”, although those involved in regenerative medicine place more emphasis on the use of stem cells to produce tissues [412]. Tissue engineering in vitro and in vivo involves the interaction of cells with a material surface. The nature of the surface can directly influence cellular response, ultimately affecting the rate and quality of new tissue formation. Initial events at the surface include the orientated adsorption of molecules from the surrounding fluid, creating a conditioned interface to which the cell responds. The gross morphology, as well as the micro-topography and chemistry of the surface, determine which molecules can adsorb and how cells will attach and align themselves. The local attachments made of the cells with their substrate determine cell shape, which, when transduced via the cytoskeleton to the nucleus, result in expression of specific phenotypes. Osteoblasts and chondrocytes are sensitive to subtle differences in surface roughness and surface chemistry. Boyan et al. [413] investigated the chondrocyte response to TiO2 of differing crystallinities, and showed that cells can discriminate between surfaces at this level as well. Cellular response also depends on the local environmental and state of maturation of the responding cells. It was mentioned that optimizing surface structure for site-specific tissue engineering is one option; modifying surfaces with biological means is another biological engineering [413].\nOne major determination of the suitability of various engineering materials for use in biological settings is the relative strength of adhesion obtained between those materials and their contacting viable phases [414]. Maximal adhesive strength and immobility are desired for orthopedic and dental implants. For example, while minimal bio-adhesion is critical to preventing unwanted thrombus formation in cardiovascular devices, plaque buildup on dental prostheses, and bacterial fouling [414]. Attention should be directed to adhesive phenomena in the oral environment, examining new surface conditioning methods for the prevention of micro-organism deposits, as well as the promotion of excellent tissue bonding to implanted prosthetic devices. Other bio-adhesive phenomena considered included those important to the safe and effective function of new cardiovascular devices [414].\nScaffold material has a two-fold function: artificial extracellular matrices (ECM) and as a spacer keeping a certain open space [415]. Furthermore, scaffold material has to be dissolved completely into the living body after auto-cell is regenerated with artificial extracellular matrices [415]. There are several important biodegradable and/or biofunctional scaffold architectures, structures and materials. They include blended-polymer scaffolds, collagen-based scaffolds, and composite scaffolds of polyhydroxybutyrate-polyhydroxyvalerate with bioactive wollastonite (CaSiO3) [416]. Using an ink-injection technique [417], a thin film (with thickness of about 0.1mm) of calcium phosphate and binding agent is injected onto the substrate to build 3-D bony-like structures [418]. Lee et al. [419] employed three-dimensional printing (3DP) technology to fabricate porous scaffolds by inkjet printing liquid binder droplets. Direct 3DP, where the final scaffold materials are utilized during the actual 3DP process, imposes several limitations on the final scaffold structure. An indirect 3DP protocol was developed, where molds are printed and the final materials are cast into the mold cavity to overcome the limitations of the direct technique. Results of SEM observations revealed highly open, well interconnected, uniform pore architecture (about 100–150 μm) [419]. Scaffold materials for bone tissue engineering often are supplemented with bone morphogenetic proteins (BMPs). Walboomers et al. [420] investigated a bovine extracellular matrix product containing native BMPs. Hollow cylindrical implants were made from titanium fiber mesh, and were implanted subcutaneously into the back of Wistar rats. It was reported that (i) after eight weeks, in two out of six loaded specimens, newly-formed bone and bone marrow-like tissues could be observed, and (ii) after 12 weeks, this had increased to five out of six loaded samples. It was, therefore, concluded that the bovine extracellular matrix product loaded in a titanium fiber mesh tube showed bone-inducing properties [420].\nElectrospinning [421] has recently emerged as a leading technique for generating biomimetic scaffolds made of synthetic and natural polymers for tissue engineering applications. Li et al. [422] compared collagen, gelatin (denatured collagen), solubilized alpha-elastin, and recombinant human tropoelastin as biopolymeric materials for fabricating tissue engineered scaffolds by electrospinning. It was reported that (i) the average diameter of gelatin and collagen fibers could be scaled down to 200–500 nm without any beads, while the alpha-elastin and tropoelastin fibers were several microns in width, and (ii) cell culture studies confirmed that the electrospun engineered protein scaffolds support attachment and growth of human embryonic palatal mesenchymal cells [422]. For fabricating meshes of collagen and/or elastin by means of electrospinning from aqueous solutions, Buttafoco et al. [423] added polyethylene oxide and NaCl to spin continuous and homogeneous fibers. It was reported that (i) upon crosslinking, polyethylene oxide and NaCl were fully leached out, and (ii) smooth muscle cells grew as a confluent layer on top of the crosslinked meshes after 14 days of culture [423].\nSurface properties of scaffolds play an important role in cell adhesion and growth. Biodegradable poly(α-hydroxy acids) have been widely used as scaffolding materials for tissue engineering; however, the lack of functional groups is a limitation. Liu et al. [424] mentioned in their studies that gelatin was successfully immobilized onto the surface of poly(α-hydroxy acids) films and porous scaffolds by an entrapment process. It was found that (i) the amount of entrapped gelatin increased with the ratio of dioxane/water in the solvent mixture used, (ii) chemical crosslinking after physical entrapment considerably increased the amount of retained gelatin on the surface of poly(α-hydroxy acids), (iii) osteoblasts were cultured on these films and scaffolds, (iv) cell numbers on the surface-modified films and scaffolds were significantly higher than those on controls 4 h and 1 day after cell seeding, (v) the osteoblasts showed higher proliferation on surface-modified scaffolds than on the control during 4 weeks of in vitro cultivation, and (vi) more collagen fibers and other cell secretions were deposited on the surface-modified scaffolds than on the control scaffolds [424].\nThere are still unique scaffold systems developed, such as the collagen-carbon nanotubes composite matrices [425], chitosan-based hyaluronan hybrid polymer fibers system [426], bioactive porous CaSiO3 scaffold structure [427], or a three-dimensional porous scaffold composed of biodegradable polyesters [428].\n\n5.8. Application of Nanotechnology to Surface Modification\nSeveral studies have suggested that materials with nanopatterned surfaces produced from various chemistries, such as metals, polymers, composites and ceramics, exhibit better osseointegration when compared to conventional materials [429–432]. Nano-patterned surfaces provide a higher effective surface area and nanocavities when compared to the conventional microrough surfaces. These properties are crucial for the initial protein adsorption that is very important in regulating the cellular interactions on the implant surface. Taylor [433] pointed out that nanotechnology offers the key to faster and remote diagnostic techniques – including new high throughput diagnostics, multi-parameter, tunable diagnostic techniques, and biochips for a variety of assays. It also enables the development of tissue-engineered medical products and artificial organs, such as heart valves, veins and arteries, liver, and skin. These can be grown from the individual’s own tissues as stem cells on a 3-D scaffold, 3-D tissue engineering extracellular matrix, or the expansion of other cell types on a suitable substrate. The applications which seem likely to be most immediately in place are external tissue grafts; dental and bone replacements; protein and gene analysis; internal tissue implants; and nanotechnology applications within in vivo testing devices and various other medical devices. Nanotechnology is applied in a variety of ways across this wide range of products. Artificial organs will demand nanoengineering to affect the chemical functionality presented at a membrane or artificial surface, and thus avoid rejection by the host. There has been much speculation and publicity about more futuristic developments such as nanorobot therapeutics, but these do not seem likely within our time horizon [433].\nThere are studies done on nanotubes. For example, Frosch et al. [434] investigated the effect of different diameters of cylindrical titanium channels on human osteoblasts. Titanium samples with continuous drill channels with various diameters (300, 400, 500, 600, and 1,000 microns) were put into osteoblast cell cultures that were isolated from 12 adult human trauma patients. It was reported that (i) within 20 days, cells grew an average of 838 μm into the drill channels with a diameter of 600 μm, and were significantly faster than in all other channels, (ii) cells produced significantly more osteocalcin messenger RNA (mRNA) in 600 μm channels than they did in 1,000 μm channels, and demonstrated the highest osteogenic differentiation, (iii) the channel diameter did not influence collagen type I production, and (iv) the highest cell density was found in 300 μm channels, suggesting that the diameter of cylindrical titanium channels has a significant effect on migration, gene expression, and mineralization of human osteoblasts [434]. Macak et al. [435] reported on the fabrication of self-organized porous oxide-nanotube layers on the biomedical titanium alloys Ti-6Al-7Nb and Ti-6Al-4V by an anodizing treatment in 1M (NH4)2SO4 electrolytes containing 0.5 wt % of NH4F. It was shown that (i) under specific anodization conditions, self-organized porous oxide structures can be grown on the alloy surface, (ii) SEM images revealed that the porous layers consist of arrays of single nanotubes with a diameter of 100 nm and a spacing of 150 nm, (iii) for the V-containing alloy, enhanced etching of the β-phase is observed, leading to selective dissolution and an inhomogeneous pore formation, and (iv) for the Nb-containing alloy an almost ideal coverage of both phases is obtained. According to XPS measurements, the tubes are a mixed oxide with an almost stoichiometric oxide composition, and can be grown to thicknesses of several hundreds of nanometers, suggesting that a simple surface treatment for Ti alloys has high potential for biomedical applications [435]. A vertically aligned nanotube array of titanium oxide was fabricated on the surface of titanium substrate by anodization. The nanotubes were then treated with NaOH solution to make them bioactive, and to induce growth of hydroxyapatite (bone-like calcium phosphate) in a simulated body fluid. It is found that (i) the presence of TiO2 nanotubes induces the growth of a “nano-inspired nanostructure”, i.e., extremely fine-scale (∼8 nm feature) nanofibers of bioactive sodium titanate structure on the top edge of the ∼15 nm thick nanotube wall, (ii) during the subsequent in vitro immersion in a simulated body fluid, the nano-scale sodium titanate, in turn, induced the nucleation and growth nano-dimensioned hydroxyapatite phase, and (iii) such TiO2 nanotube arrays and associated nanostructures can be useful as a well-adhered bioactive surface layer on Ti implant metals for orthopedic and dental implants, as well as for photocatalysts and other sensor applications [436].\nWebster et al. have suggested that enhanced vitronectin adsorption, conformation and bioactivity are the major reason for increased osteoblast adhesion on nanophase alumina [437]. In the last years, a new method has been described to fabricate nanotubular structures on titanium surfaces. These titania nanotubes can be produced by a variety of methods including electrochemical deposition, sol-gel method, hydrothermal processes and anodic oxidation [438–440]. Using this novel approach, several studies showed that the presence of the nanotube structure on a titanium surface induced a significant increase in the action of osteoblastic cells compared to those grown on flat titanium surfaces [441,442]. Nanotubular TiO2 layer produced using anodization has an amorphous crystal structure and it has been shown that using heat-treatment it can be transformed into anatase to improve cellular interactions [441]. In this study, a sintering protocol at 450 °C for 2 h was used to perform a crystal phase transformation of nanotubes. Significantly higher cell proliferation rates and better cellular morphologies were observed on anatase nanotubular surfaces after 7 days of culture, as shown in the literature [442]. Park et al. [443] produced nanotubular surfaces having pore diameter of 15, 20, 30, 50, 70 and 100 nm without heat treatment and documented that on nanotubular surfaces above 50 nm, the cell attachment and spreading was significantly decreased, thereby causing an increased programmed cell death. They only showed better cell proliferation and matrix mineralization results on nanotubes having 15 nm pore diameter. Whereas, in another study, Oh et al. [444] performed heat treatment following anodization and showed that anatase nanotubes having larger pore diameter (70 to 100 nm) MSCs elongate better and undergo selective differentiation into osteoblast-like cells compared to small nanotubes. In the present study, nanotubes having pore diameter of 70–100 nm were produced and impaired cellular functions on non heat-treated amorphous nanotubular surfaces were observed. However, on anatase nanotubular surfaces, significantly increased cell proliferation values were recorded after 7 days of culture, as shown in the literature [441,442,444]. Residual fluorine within the amorphous nanotubes following anodization might be the factor for this decreased cellular density, as stated in the literature [441]. Therefore, this study was able to show that heat treatment is essential for the production of nanotubular implant surfaces since it provides a more ideal oxide crystal structure for the spreading and proliferation of the cells.\nBeside microtopographical features, surface wettability and surface free energy are also important parameters influencing cell attachment, proliferation and differentiation [445]. Bauer et al. [446] cultured rat mesenchymal stem cells on nanotubular titanium surfaces having different wettability characteristics and found an increased cell attachment on super-hydrophobic surfaces compared with super-hydrophilic ones. In the present study, samples having different wettability profiles were obtained following roughening, anodization and heat treatment procedures. After anodization and heat treatment, water contact angles decreased gradually. However, due to the changes in surface chemistries following treatments, no correlation was found between surface wettability and cellular functions. Further studies are needed to evaluate the effect of hydrophilicity and surface chemistry of titania nanotubes on protein adsorption and cell responses.\nThree types of bioactive polymethylmethacrylate (PMMA)-based bone cement containing nano-sized titania (TiO2) particles were prepared, and their mechanical properties and osteoconductivity are evaluated by Goto et al. [447]. The three types of bioactive bone cement were un-silanized TiO2, 50 wt%, silanized TiO2 50 wt%, and 60 wt% mixed to PMMA. Commercially available PMMA cement was used as a control. The cements were inserted into rat tibiae and allowed to solidify in situ. After 6 and 12 weeks, tibiae were removed for evaluation of osteoconductivity. It was reported that (i) bone cements using silanized TiO2 were directly apposed to bone, while un-silanized TiO2 cement and PMMA control were not, (ii) the osteoconduction of cement with 60 wt% of silanized TiO2 was significantly better than that of the other cements at each time interval, and (iii) the compressive strength of cement with 60 wt% of silanized TiO2 was equivalent to that of PMMA, indicating that cement with 60 wt% of TiO2 was a promising material for use as a bone substitute [447]. Since it is essential for the gap between the hydroxyapatite coated titanium and juxtaposed bone to be filled out with regenerated bone, promoting the functions of bone-forming cells is desired. In order to improve orthopedic implant performance, Sato et al. [448] synthesized nanocrystalline hydroxyapatite (HA) powders to coat titanium through a wet chemical process. The precipitated powders were either sintered at 1,100 °C for 1 h in order to produce microcrystalline size HA, or were treated hydrothermally at 200 °C for 20 h to produce nanocrystalline HA. These powders were then deposited onto titanium by a room temperature process. It was reported that (i) the chemical and physical properties of the original HA powders were retained when coated on titanium by the room temperature process, (ii) osteoblast adhesion increased on the nanocrystalline HA coatings compared to traditionally used plasma-sprayed HA coatings, (iii) greater amounts of calcium deposition by osteoblasts cultured on Y-doped nanocrystalline HA coatings were observed [448]. With a wide variety of applications, nanotechnology has attracted the attention of researchers as well as regulators and industrialists, including nanodrugs and drug delivery, prostheses and implants, and diagnostics and screening technologies. We can take advantages of availability of ultra-fine nanomicrostructures of solid metals, alloys, powder, fibers, or ceramics to fabricate superplastically formed products.\nIt is indispensable here to mention the minimally invasive dentistry (MID) and minimally invasive surgery (MIS). The MIS concept has been created to allow new thinking and a new approach to dentistry where restoration of a tooth becomes the last treatment decision rather than first consideration as at present. It provides a practical approach to caries preventive measures based on the notion of demineralization and remineralization in a micro-phase in order to retain healthy teeth. The medical model of MID is characterized by (1) reduction in cariogenic bacteria, (2) preventive measures, (3) remineralization of early enamel lesions, (4) minimum surgical intervention of cavitated lesions, and (5) repair of defective restorations [449]. At the same time, it is mentioned that MIS has several advantages: (1) since the surgical area is narrower, damage on surrounding soft tissue can be minimized, (2) post-operation pain can be minimized, (3) hospital time can be shortened, and (4) early rehabilitation can be initiated [450]. These MIs in both dentistry and medicine inevitably require precisely manufactured prostheses in micro-scale or even nano-scale. It is anticipated that the MI-based technologies, as well as MI-oriented technologies, will be advanced in the near future.\n\n5.9. Bioengineering and Biomaterial-Integrated Implant System\nWith the aforementioned supportive technologies, surfaces of dental and orthopedic implants have been remarkably advanced. These applications can include not only ordinal implant system but also miniaturized implants, as well as customized implants. Dental implant therapy has been one of the most significant advances in dentistry in the past 25 years. The computer and medical worlds are both working hard to develop smaller and smaller components. Using a precise, controlled, minimally invasive surgical (MIS) technique, the mini dental implants (MDI) are placed into the jawbone. The heads of the implants protrude from the gum tissue and provide a strong, solid foundation for securing the dentures. It is a one-step procedure that involves minimally invasive surgery, no sutures, and none of the typical months of healing. Advantages associated with the MDI are (1) it can provide immediate stabilization of a dental prosthetic appliance after a minimally invasive procedure. (2) It can be used in cases where traditional implants are impractical, or when a different type of anchorage system is needed. (3) Healing time required for mini-implant placement is typically shorter than that associated with conventional 2-stage implant placement and the accompanying aggressive surgical procedure. According the clinical reports, a biometric analysis of 1,029 MDI min-implants, five months to eight years in vivo showed that the MDI mini-implant system can be implemented for long-term prosthesis stabilization, and delivers a consistent level of implant success [451].\nIn addition to the aforementioned miniature implants, an immediate loading, as well as customized implants, have been receiving attention recently. Conventionally, a dental implant patient is required to have two-stages of treatment consisting of two dental appointments five to six months apart. Recently, a single stage treatment has received attention. Placing an implant immediately or shortly after tooth extraction offers several advantages for the patient as well as for the clinician. These advantages include shorter treatment time, less bone sorption, fewer surgical sessions, easier definition of the implant position, and better opportunities for osseointegration because of the healing potential of the fresh extraction site [452–455]. Titanium bar (particularly the portions in direct contact to connective tissue and bony tissue) is machined to have the exact shape of the root portion of the extracted tooth of the patient. The expected outcome of this method is a perfect mechanical retention, and therefore an ideal osseointegration can be achieved. This is called a custom (or customized) implant, which is fabricated by the electro-discharge machining (EDM) technique.\nPrefabricated dental implant can be further machined to replicate the extracted tooth and machined implant (whose shape is exactly same as the patient’s extracted tooth) can be prepared within a relatively short operation time and can be placed within one hour to the patient. Since the root shape of the placed implant is exactly same as the extracted tooth’s root form, the follow-up reaction including osseointegration is expected shorter and placed implant’s retention force can be achieved within a relatively short time.\nElectron-beam machining (EBM) is a machining process where high-velocity electrons (in the range of 50 to 200 kV to accelerate electrons to 200,000 km/s) are directed toward a work piece, creating heat and vaporizing the material. Electromagnetic lenses are used to direct the electron beam, by means of deflection, into a vacuum. The electrons strike the top layer of the work piece, removing material, and then become trapped in some layer beneath the surface. Applications of this process are annealing, metal removal, and welding. EBM can be used for very accurate cutting of a wide variety of metals. Surface finish is better and kerf width is narrower than those for other thermal cutting processes. The process is similar to laser-beam machining, but because EBM requires a vacuum, it is not used as frequently as laser-beam machining [456]. In addition to the immediate placement of dental implants, another concept has been introduced. Techniques such as stereoscopic lithography and computer-assisted design and manufacture (CAD/CAM) have been successfully used with computer-numerized control milling to manufacture customized titanium implants for single-stage reconstruction of the maxilla, hemimandible, and dentition without the use of composite flap over after the removal of tumors [457]. Nishimura et al. [458] applied this concept to dental implants to fabricate the individual and splinted customized abutments for all restoration of implants in partially edentulous patients. It was claimed that complicated clinical problems such as angulation, alignment, and position can be solved. However, with this technique, the peri-implant soft tissues are allowed to heal 2 to 3 weeks, so that at least two dental appointments are required.\nMany oral implant companies (about 25 companies are currently marketing 100 different dental implant systems) have recently launched new products with claimed unique, and sometimes bioactive surfaces [459,460]. The focus has shifted from surface roughness to surface chemistry and a combination of chemical manipulations on the porous structure. To properly explain the claims for new surfaces, it is essential to summarize current opinions on bone anchorage, with emphasis on the potentials for biochemical bonding. There were two ways of implant anchorage or retention: mechanical and bioactive [459,460].\nRecent research has further redefined the retention means of dental implants into the terminology of osseointegration versus biointegration. When examining the interface at a higher magnification level, Sundgren et al. [30] showed that unimplanted Ti surfaces have a surface oxide (TiO2) with a thickness of about 35 nm. During an implantation period of eight years, the thickness of this layer was reported to increase by a factor of 10. Furthermore, calcium, phosphorous, and carbon were identified as components of the oxide layer, with the phosphorous strongly bound to oxygen, indicating the presence of phosphorous groups in the metal oxide layer. Many retrospective studies on retrieved implants, as well as clinical reports, confirm the aforementioned important evidence (1) surface titanium oxide film grows during the implantation period, and (2) calcium, phosphorous, carbon, hydroxyl ions, proteins, etc. are incorporated in an ever-growing surface oxide even inside the human biological environments [460,461]. Numerous in vitro studies, (e.g., [461]) on treated or untreated titanium surfaces were covered and to some extent were incorporated with Ca and P ions when such surfaces were immersed in SBF (simulated body fluid). Additionally, we know that bone and blood cells are rugophilia, hence in order not only to accommodate for the bone growth, but also to facilitate such cells adhesion and spreading, titanium surfaces need to be textured to accomplish and show appropriate roughness [462]. Furthermore, gradient functional concept (GFC) on materials and structures has been receiving special attention not only in industrial applications, but in dental as well as medical fields [462]. Particularly, when such structures and concepts are about to be applied to implants, its importance becomes more clinically crucial. For example, the majority of implant mass should be strong and tough, so that occlusal force can be smoothly transferred from the placed implant to the receiving hard tissue [462]. However, the surface needs to be engineered to exhibit some extent of roughness. From such macro-structural changes from bulk core to the porous case, again the structural integrity should be maintained. The GFC can also be applied for the purpose of having a chemical (compositional) gradient. Ca-, P-enrichment is not needed in the interior materials of the implants. Some other modifications related to chemical dressing or conditioning can also be utilized for achieving gradient functionality on chemical alternations on surfaces as well as near-surface zones [462].\n\n5.10. Technology-Integrated Implant Systems\nAny new type of implant (not only dental but also orthopedic applications) should possess a gradual function of mechanical and biological behaviors, so that mechanical compatibility and biological compatibility can be realized with s single implant system [462]. Since microtextured Ti surfaces [67,463,464] and/or porous Ti surfaces [465–467] promote fibroblast apposition and bone ingrowth, the extreme left side representing the solid Ti implant body should have gradually increased internal porosities toward to the case side which is in contact with vital hard/soft tissue. Accordingly, mechanical strength of this implant system decreases gradually from core to case, whereas biological activity increases from core to case side. Therefore, the mechanical compatibility can be completely achieved. Porosity-controlled surface zones can be fabricated by an electrochemical technique [468], polymeric sponge replication method [65], powder metallurgy technique, superplastic diffusion bonding method [469], or foamed metal structure technique [470].\nOnce the Ti implant is placed in hard tissue, TiO2 grows and increases its thickness [30,41,471–478], due to more oxygen availability inside the body fluid, as well as co-existence of superoxidant. It is very important to mention here that Ti is not in contact with the biological environment, but rather there is a gradual transition from the bulk Ti material, stoichiometric oxide (i.e., TiO2), hydrated polarized oxide, adsorbed lipoproteins and glycolipids, portroglycnas, collagen filaments and bundles to cells [40]. Such gradient functional structure can be also fabricated in CpTi and microtextured polyethylene terephthalate (PET) system [479]. In addition, a gradient structural system of Ti and TiN was developed [480]. During HA coating, a gradient functional layer was successfully fabricated [214]. To promote these gradient functional (GF) and gradient structural (GS) transitions, there are many in vivo, as well as in vitro, evidences indicting that surface titanium oxide is incorporated with mineral ions, water and other constituents of biofluids [30,38,39,481]. Since a surface layer of TiO2 is negatively charged, the calcium ion attachment can be easily achieved [33,482]. Retrieved Ti implants showed that surface TiO2 was incorporated with Ca and P ions [483], while in vitro treatment of TiO2 in extracellular fluids or simulated body fluid (SBF) for prolonged periods of incubation time resulted in the incorporation of Ca, P, and S ions into TiO2 [30,38–41,218,471,481,484]. Without prolonged treatment, there are several methods proposed to relatively short-time incubation for incorporation of Ca and P ions. For example, TiO2 can be electrochemically treated in an electrolyte of a mixture of calcium acetate monohydrate and calcium glycerophosphate [145]. As a result of incorporation of Ca and P ions, bone-like hydroxyapatite can be formed in macro-scale [485] or nano-dimension [448]. Again for reducing the incubation time, bone-like hydroxyapatite crystals can be formed by treating the TiO2 surface with water and hydrogen plasma immersion ion implantation, followed by immersion in SBF [146], or by treating in hydrogen peroxide followed by SBF immersion [147], or immersion in SBF while treating the TiO2 surface with micro-arc oxidation and irradiation with UV light [171]. It is also known that P ions can be incorporated into TiO2 while it is immersed in the human serum [40].\nBony growth extreme surface zones should have a same roughness as the roughness of receiving hard tissue through micro-porous texturing techniques. This area can be structured using nanotube concepts [434–436]. Because this zone responds strongly to osseointegration, the structure, as well as the chemistry, should accommodate favorable osteoinductive reactions. Bone morphogenetic protein [385,486,487], and nano-apatite can be coated [488]. The zone may be treated by femtosecond laser machining [404] to build a micro-scale 3D scaffold which is structured inside the macro-porosities. Such scaffold can be made of biodegradable material (e.g., chitosan), which is incorporated with protein, Ca, P, apatite particles or other species possessing bone growth factors [483].\n"}